USPTO Patents Application 10687850

J

Europaisches Patentamt
European Patent Office
Office europ^en des brevets

© Publication number:

Piillilllllllii

0 640 832 A2

0 Application number: 94113523^
<§) Date of filing: 30.08.94

EUROPEAN PATENT APPLICATION

® Int. CIA G01N 27/327, G01N 33/543

® Priority: 30.0a93 US 112926

@ Applicant: Hughes Aircraft Company

® Date of publication of application:

7200 Hughes Terrace,

P.O. Box 80028

01.03.95 Bulletin 95/09

Los Angeles,

@ Designated Contracting States:

Califomia 90080-0028 (US)

@ Inventor: Kindler, Andrew

DE ES PR GB IT

616 Plymouth Road

San Marino,

California 91108 (US)

@ Representative: Ross, W

KUHNEN, WACKER 8t PARTNERS,

Alois-Steinecker-Strasse 22

D-85354 Freising (DE)

@ Electrochemical immunosensor system.

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® A electrochemical immunosensor system is pro-
vided which uses electrical signals to measure bind-
ing events. The system includes an immunosensor
having a sensing electrode in contact with a fluid
containing an analyte concentration. A quantity of
antibodies or other binding agent is adsorbed on or
otherwise affixed to the electrode surface such that a
portion of the antibodies of the binding agent binds a
portion of the analyte to form complexes on the
electrode surface. Signal generating means develop
an electrical signal at the sensing electrode such
that a response current is produced through the
sensing electrode. The response current has mea-
surable signal characteristics which are dependent
upon the number of complexes formed, and there-
fore the analyte concentration within the fluid.

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BACKGROUND OF THE INVENTION

1 . Reld of the Invention

The present invention relates generally to im-
munologic measurement systems, in which anti-
body binding is used to measure substance con-
centration levels. More particularly, the present in-
vention relates to an electrochemical immunosen-
sor in which voltametric signals are used to convert
immunosensor binding events to a measurable
electrical quantity.

2. Description of Related Art

Immunosensors are a subset of the class of
biological measurement instruments commonly
known as biosensors. Biosensors typically consist
of probes containing biological recognition mol-
ecules. The recognition molecules respond to the
presence of certain substances, and the response
is then converted to a measurable quantity. Biosen-
sors thus provide an indication of substance con-
centration without the use of complicated labora-
tory procedures. An immunosensor is a particular
type of biosensor in which an antibody serves as
the biological recognition molecule. Antibodies are
produced in the human body to inactivate foreign
substances by irreversibly combining with or bind-
ing the substance to form a complex. An almost
unlimited variety of antibodies are produced, each
specific to a particular substance. The term "anti-
gen" as used herein refers to any substance that
can cause the immune system to form antibodies.
The term "analyte" will be used to refer to any
substance, including an antigen, which binds with
any other substance to form a complex. Sub-
stances which bind analytes will be referred to as
"binding agents". An antibody is one particular
type of binding agent. Other binding agents include
receptors and affinity-binding molecules.

The highly sensitive and selective nature of
antibody binding permits the immunosensor to ac-
curately detect minute analyte concentration levels.
Immunsensors are therefore useful in medical ap-
plications involving the detection of hormones, il-
legal drugs or other trace constituents present in
blood and urine. Furthermore, since antibodies can
be selected to bind to a wide variety of different
analytes, the Immunosensor has the potential for
widespread application to many non-medical uses
such as industrial and food processing quality con-
trol, monitoring of environmental pollutants and de-
tection of materials used in chemical and biological
wartare.

Despite these advantages and potentially wide-
spread applications, immunosensors are not com-
mercially available at the present time. A significant

obstacle has been the problem of accurately con-
verting the antibody recognition event to a measur-
able signal. As a result, more complex measure-
ment techniques are currently used in place of

5 immunosensors. One such group of techniques in-
volves multi-step analytic laboratory procedures
known as immunoassays. Several exemplary im-
munoassays are described in E. Harlow and D.
Lane, Antibodies: a Laboratory Manual, pp. 553-

10 612. In applications requiring a highly accurate
measurement, the immunoassay techniques used
typically require highly trained technicians and
costly special equipment. The radioimmunoassay
("RIA") is a good example of this type of highly

75 accurate immunoassay. Measurements are there-
fore typically performed at a fully equipped central
laboratory rather than at remote user locations such
as a medical office or a factory. The complex
conversion process results in a significant delay

20 before availability of measurement results.

In applications in which a lower degree of ac-
curacy is acceptable, immunoassay techniques
which make use of subjective visual examination
are typically used. Such techniques provide a con-

25 venient but less accurate screen in variety of situ-
ations. The subjective tests are often of the
"dipstick" type. A sensitized material is placed in
the analyte, and a color change on the surface of
the material occurs. The pregnancy tests available

30 in most supermarkets are usually of this type. This
type of test generally belongs to a class of im-
munoassays known as enzyme linked immunosor-
bent assay ("ELISA"). Although ELISA techniques
can provide fast results, they are intrinsically less

35 accurate than radioimmunoassay techniques.

There are additional problems with particular
immunoassay techniques. For example, one of the
most sensitive of the immunoassays is the two-step
sandwich type immunoradiometric assay ("IRMA")

40 in which the complexation of an antibody with an
antigen is detected by measuring the amount of
radioactivity emitted by a radiolabeled antibody.
The use of radioactive labels contributes to sen-
sitivity, but exposes laboratory personnel to a sig-

45 nificant safety hazard. Moreover, radioactive com-
pounds have a limited shelf life. Other immunoas-
say techniques, such as fluorescence polarization
immunoassay ("FPIA") and ELISA. are safer for
laboratory personnel and more stable with regard

60 to shelf life, but are generally less sensitive than
the IRMA.

Another alternative to electrochemical im-
munosensors uses light scattering to convert the
antibody binding response to a measurable elec-
55 tricat signal. One such technique uses a spectrom-
eter to measure the variation in the spectrum of
laser light passing through a solution containing
antibodies and antigens. The antibody-antigen

2

EP 0 640 832 A2

binding response alters the spectral characteristics
of the laser in accordance with the antigen con-
centration level. U.S. Patent No. 4.725,149 is a
variation on this general technique in which laser
light is passed through a solution containing anti-
gens and magnetic particles coated with anti-
bodies. The magnetic particles are rotated at a
particular frequency by signals applied to coils
surrounding the solution such that antibody-antigen
binding events produce a measurable variation in
the scattered light. In U.S. Patent No. 4,762.413 a
photodetector measures the power spectral density
of fluctuations in scattered light intensity* and a
mean power spectral density value taken over sev-
eral measurements is used to indicate antigen con-
centration.

Light scattering conversion techniques such as
these, however, typically utilize expensive specially
designed equipment which would tend to limit use
at remote sites. In addition, the techniques gen-
erally require combining the measurement sample
with a buffer solution containing a quantity of par-
ticles. The antibodies which will bind to the anti-
gens in the sample are fixed to the surfaces of the
particles. Thus, the light scattering techniques are
not readily adaptable to those situations in which a
sample cannot be easily removed and combined
with the particle solution. Furthermore, since the
recognition molecules are placed within the particle
solution rather than on a probe, the light scattering
techniques do not share the advantages of conve-
nience and simplicity commonly associated with
biosensors.

Attempts have been made to simplify optical
Immunosensor detection systems such as those
discussed above by attaching antibodies or other
binding agents to a deformable organic polymer
film. The polymer film absorbs green light and
fluoresces strongly in the orange part of the spec-
trum. When an antibody attached to a surface of
the film combines with an analyte, the polymer film
is perturbed, causing a detectable decrease in the
fluorescence of the film at the point of the com-
bination. The decrease in fluorescence is propor-
tional to the number of analyte molecules bound to
antibodies on the surface of the film. The amount
of light reflected from the film is measured to
determine the analyte concentration in a particular
sample. See "Signal Transduction on Film" in
Bioventure View, March 1992. Although the poly-
mer film approach may decrease the complexity of
the optical immunosensor, it presents additional
problems. One such problem is the inability to
place a sufficient quantity of fcMnding agents at
particular points on the film.

As is apparent from the above, there presently
is a need for a simple and inexpensive im-
munosensor system in which antibody recognition

events are rapidly and accurately converted to
readily monitored electrical signals using standard
electronic test equipment. The immunosensor sys-
tem should provide highly sensitive, selective and

5 repeatable measurements without the problems as-
sociated with immunoassay or light scattering tech-
niques. Furthermore, the immunosensor and re-
lated equipment should be easily adapted to the
specific requirements of a variety of different uses.

10 including but not limited to medical, industrial, envi-
ronmental and military applications.

SUMMARY OF THE INVENTION

T5 The present invention uses voltametric elec-
trochemical analysis techniques to accurately con-
vert immunosensor binding events to easily mea-
sured quantities. In the electrochemical im-
munosensor system of the present invention, at

20 least one sensing electrode is provided in contact
with a fluid containing an analyte concentration to
be measured. The sensing electrode has a surface
on which a quantity of a binding agent is present. A
portion of the binding agent binds a portion of the

25 analyte within the fluid to form a number of com-
plexes. Signal generating means develop an elec-
trical signal at the electrode, producing a response
current signal which is monitored by signal moni-
toring means. The response current has at least

30 one signal characteristic dependent upon the num-
ber of complexes formed, and is therefore indica-
tive of the analyte concentration level within the
fluid.

In accordance with one aspect of the present

35 invention, a method for measuring analyte con-
centrations is provided in which ac voltametric sig-
nal is applied to the equipment connected to the
immunosensor electrode in order to generate an
appropriate response current Signal characteristics

40 of the response current second harmonic provide
an accurate indication of analyte concentration.

In accordance with another aspect of the
present invention, a method is provided in which
the fluid contains a quantity of indicator ions and

45 wherein an electrical signal is applied to the elec-
trode to generate an ionic response current. The
magnitude of the response current varies as a
function of the number of antibodies on the elec-
trode surface which have bound analytes in the

so fluid, which is in turn a function of the analyte
concentration.

As a feature of the present invention, the elec-
trochemical immunosensor uses voltametric mea-
surement techniques and readily available test

55 equipment to accurately convert immunosensor
antibody binding responses to measurable elec-
trical signals.

3

EP 0 640 832 A2

As an additional feature of the present inven-
tion, the high sensitivity and selectivity of an im-
munosensor is achieved without the cost and com-
plexity of the present immunoassay or light scatter-
ing conversion techniques. The sensor or electrode
portion of the electrochemical immunosensor of the
present invention can be inexpensively manufac-
tured and therefore discarded after each use. A
disposable immunosensor avoids the problems and
additional cost associated with cleaning and resur-
facing a used sensor.

As a further feature of the present invention,
the techniques disclosed can produce inexpensive
yet accurate results in many different applications
using the same standard set of test equipment and
measurement techniques. The electrochemical im-
munosensor of the present invention therefore per-
mits widespread commercial immunosensor use,
resulting in significant benefit to users in many
fields. For example, the simple conversion tech-
niques of the present invention will permit remote
medical clinics to perform rapid and inexpensive
analysis of a variety of constituents, including
blood, urine and saliva.

The above-discussed features and attendant
advantages of the present invention will become
better understood by reference to the following
detailed description of the preferred embodiment
and the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic representation of an
exemplary preferred embodiment of the electro-
chemical immunosensor system of the present in-
vention.

FIG. 2 is a schematic representation of a sec-
ond preferred embodiment of the electrochemical
immunosensor system of the present invention.

DETAILED DESCRIPTION OF THE PREFERRED
EMBODIMENT

The present invention provides a simple and
efficient means of converting immunosensor bind-
ing events to readily measurable quantities using
voltametric electrochemical analysis techniques.
One such voltametric technique, disclosed in U.S.
Patent No. 4,631,116, and assigned to the present
assignee, uses voltametric signals to produce ac
current spectra which vary as a result of changes
in the concentration of plating bath trace constitu-
ents. The contents of U.S. Patent No. 4.631,116 are
hereby expressly incorporated by reference. Al-
though the following detailed description is primar-
ily directed to using the present invention In con-
junction with antibody-antigen binding and exem-
plary ac voltametric techniques, this is by way of

example and not limitation. It should be understood
that the system and methods described can be
used to detect the concentration of any analyte
which binds with any binding agent to form a

5 complex. Many other ac and dc electrical signals
or voltametric analytic techniques could also be
used to detect binding events in accordance with
the present invention.

The schematic diagram of FIG. 1 shows an

10 exemplary preferred embodiment of the electro-
chemical immunosensor system of the present In-
vention. The fluid to be measured Is located within
a sensor 9. The sensor 9 can be submerged within
a fluid container or a sample of the fluid may be

75 placed within the sensor. In the case of measure-
ments of bodily fluids, the sensor could be inserted
Into the body or fluid samples could be removed
from the body and placed within the sensor 9 in
diluted or undiluted form.

20 The sensor 9 contains a sensing electrode 10
which Is preferably constructed of an Inert material
such as gold or platinum. Gold is preferred in
many applications since it is both inert and known
to adsorb antibodies and other proteins very effi-

25 ciently. A sufficient quantity of antibodies can thus
be adsorbed and retained to prevent significant
loss of antibodies back to the fluid during voltage
excursions into the desorption range of the vol-
tametric signal potential. Platinum has less ability

30 to adsorb antibodies but may be preferred to gold
in those applications which do not require as high a
rate of antibody adsorption at the electrode. Other
inert materials capable of adsorbing antibodies or
other binding agents could also be used.

35 The sensing electrode 10 is connected to port
25 of a potentiostat 8 via line 28. The sensor 9 also
contains a counter electrode 12 and a reference
electrode 11. All system measurements are taken
relative to the reference electrode 11. The refer-

40 ence electrode can be a standard calomel elec-
trode or any other suitable reference electrode.
Reference electrode 11 and counter electrode 12
are connected to ports 26, 27 of potentiostat 8 via
lines 29, 30, respectively. The three-electrode sen-

45 sor 9 with electrodes 10, 11 and 12 Is a sensor
design typically used in conjunction with vol-
tametric techniques. It should be understood, how-
ever, that alternative electrode arrangements may
also be used with the method of the present inven-

60 tion.

The basic operation of the equipment of FIG. 1
will now be described. A function or waveform
generator 5 provides an output 13 which is an ac
electrical signal of appropriate waveform, frequency
55 and amplitude. The ac signal is applied to the
external Input 23 of potentiostat 8 and to the refer-
ence input 16 of a lock-iri amplifier 6. The ac signal
applied to external input 23 of potentiostat 8 may

4

7 EP 0 640 832 A2 8

be superimposed upon an appropriate dc voltage
sweep signal generated by a sweep signal gener-
ator (not shown) housed within the potentiostat
enclosure. The sweep signal generator is a sepa-
rate device but is contained within the housing
enclosure of potentiostat 8 for convenience. Alter-
natively, an external sweep signal generator lo-
cated outside the potentiostat enclosure could be
used to provide the dc sweep signal. The ac signal
super-imposed on the dc sweep signal is one type
of exemplary ac voltametric signal. The ac signal
alone is another type of exemplary voltametric sig-
nal. When the exemplary voltametric signal Is ap-
plied to potentiostat 8, a signal substantially equiv-
alent to the applied voltametric signal develops
between sensing electrode 10 and reference elec-
trode 1 1 . Potentiostat 8 insures that the developed
signal voltage is maintained at the correct value by
appropriate application of cun-ent to sensor 9.

In the electrochemical immunosensor of the
present invention the antibodies or other binding
agents are preferably applied to the electrode sur-
face in advance, before the electrode comes into
contact with the fluid to be measured. The anti-
bodies may be applied to the electrode surface
during the sensor manufacturing process by ad-
sorption or any other suitable technique. Non-elec-
trochemical electrode pretreatment techniques may
be used to remove contaminants and otherwise
condition the electrode surface for adsorption or
attachment of antibodies. Although the electrode
material will generally adsorb antibodies without the
application of an adsorption signal or potential to
the electrode, an adsorption signal may be applied
to speed up the adsorption process. An electro-
chemical conditioning signal may be used prior to
measurement to stabilize an electrode which al-
ready has antibodies adsorbed or otherwise at-
tached to Its surface. Potentiostat 8 or waveform
generator 5 can serve to generate an appropriate
electrode conditioning signal. The conditioning sig-
nal voltage should be carefully monitored to avoid
removing or destroying previously attached anti-
bodies.

As discussed above, a signal substantially
equivalent to the exemplary voltametric signal ap-
plied to the potentiostat 8 is developed between
sensing electrode 10 and reference electrode 11
within sensor 9 by application of a current to elec-
trodes 10 and 12. The cun*ent is supplied by ports
25 and 27 of the potentiostat via lines 28 and 30
respectively. The current responsible for generating
the developed signal has signal characteristics
which vary depending upon the electrochemical
processes occun-ing at the surface of tiie sending
electrode 1 0. This current will be referred to herein
as the response current It is the cunrent generated
in response to the applied voltametric signal. Since

the electrochemical processes at the sensing elec-
trode surface are a function of antibody-antigen
binding, the response current is indicative of anti-
gen concentration in a manner to be further dis-

5 cussed below.

The response current passes back through
potentiostat 8, from the potentiostat output 24 to
the signal input 17 of lock-In amplifier 8. The lock-
in amplifier 6 separates out a desired ac harmonic

10 from the combined ac and dc components of the
response current signal and resolves the ac har-
monic into an In-phase and a quadrature compo-
nent. Although either the first or second ac signal
harmonic will be most commonly monitored, other

75 harmonics may provide better spectral resolution in
a particular application.

The in-phase component of the ac portion of
the response current Is then passed from In-phase
output 18 of lock-in amplifier 6 to a display signal

20 input 31 of strip chart recorder 7. Similarly, the
quadrature component is passed from quadrature
output 19 of iock-in amplifier 6 to a second display
signal input 32 of strip chart recorder 7. The strip
chart recorder displays ttie in-phase and the

25 quadrature components of the selected harmonic of
the ac response current as a function of the dc
sweep portion of the response cunrent. The quadra-
ture component will include signal peaks indicative
of changes in the electrode double layer capaci-

30 tance. The reference signal supplied from the
waveform generator 5 to Input 16 of lock-in am-
plifier 6 is related to the response current signal
and permits lock-in amplifier 6 to make accurate
relative measurements. Alternative signal monltor-

35 ing means Include digital data acquisition systems,
oscilloscopes or any other equipment suitable for
displaying or measuring tiie response current sig-
nal. The response current displays represent
unique spectra which indicate the analyte concen-

40 tration in the fluid within sensor 9.

The particular equipment used in the exem-
plary system of FIG. 1 Included a Wavetek Model
188 waveform generator, a PAR 273 potentiostat,
and a PAR 5208 lock-In amplifier. The Wavetek

45 waveform generator is available from Wavetek, San
Diego, California and the PAR equipment is avail-
able from Princeton Applied Research, Princeton,
New Jersey.

The following exemplary methods apply the

50 above described electrochemical immunosensor
system to the measurement of antibody binding
events. In general, a quantity of antitxxJies are
present on the surface of sensing electrode 10.
Although, monoclonal antibodies are preferred,

55 polyclonal or other types of antibodies could be
used. A portion of the antibodies on the electrode
surface bind a portion of the antigens in the fluid to
form complexes on the electrode surface. It should

5

9

EP 0 640 832 A2

10

again be emphasized that the use of antibodies In
this description is exemplary only, and that other
binding agents may be used in place of antibodies.

The entire sensor 9 or individual electrodes
therein may be disposable. A disposable sensor or
electrode is desirable because the strength of anti-
body-antigen binding typically makes it difficult to
separate antigens from the antibodies on the elec-
trode surface after they have combined to fonr) a
complex. A disposable sensor or electrode thus
avoids the problems and additional cost associated
with cleaning and resurfacing a used sensor. The
sensor or electrode can be inexpensively manufac-
tured and supplied in user-ready form. The sensor
or electrode can then simply be detached from the
remainder of the equipment shown in FIG. 1 and
discarded after each measurement.

The first method of converting a binding event
to a measurable electrical quantity is based upon
the detection of changes in the electrochemical
double layer capacitance at the sensing electrode
10. This method involves the application of a vol-
tametric signal consisting of an ac signal superim-
posed on a dc sweep signal to potentiostat 8. The
signal voltage developed at sensing electrode 10
mimics that of the voltametric signal applied to the
potentiostat 8. The developed signal Is therefore
also an ac signal superimposed on a dc sweep.
The dc sweep varies linearly over a potential range
at a suitable sweep rate.

The antibodies on the electrode surface will
bind the corresponding antigen in the fluid irrevers-
ibly and with great specificity. At a certain sweep
voltage, there will be a tendency for antibodies on
the sensing electrode surface to desorb. The com-
plex which is formed as a result of the antibody-
antigen binding event will tend to have a different
net charge than the antibody alone due to the
charge of the antigen. This is particularly true of
most proteins, Including hormones. Even if the
antigen Itself is neutral, a charge redistribution with-
in the complex may occur. These charge variations
will shift the voltage at which the complex will
adsorb and desorb at the electrode relative to an
unbound antit>ody.

At the transition point between absorption and
desorption for a particular type of antibody-antigen
complex, the ac signal superimposed on the dc
sweep will cause the combined ac and dc vol-
tametric signal voltage to oscillate about the aver-
age voltage required for transition between the
absorbed and desort)ed states. The oscillation be-
tween the adsorbed and desorbed states causes
changes in the response current which can be
monitored via lock-in amplifier 6 and strip chart
recorder 7 as previously described. The second
harmonic of the quadrature current is proportional
to the antigen concentration within the fluid. The

magnitude or peak of this response current oscilla-
tion is proportional to the antigen concentration
within the fluid. The second harmonic measure-
ment is preferably monitored because the constant

5 double layer capacitance portion of the response
current is thereby effectively filtered out, while the
portion of the response current reflecting changes
in double layer capacitance is displayed on re-
corder 7. Very small changes in double layer ca-

10 pacitance can be detected, corresponding to trace
levels of antibody-antigen complexes.

The sensitivity of sensing electrode double lay-
er capacitance permits measurement of the antigen
concentrations as low as 10"^^ moles/liter. For ex-

75 ample, the method can be used to detect certain
blood and urine components such as hormones,
typically found In concentrations of about 10"^ to
10~"^2 moles/liter, and illegal drug residues, at about
10-2 to 10-9 moles/liter.

20 The high measurement sensitivity of the
present invention is a result of a numt»er of factors.
Oie is the fact that the antibodies at tiie electrode
bind so strongly to the analyte that they tend to
concentrate the analyte at the electrode. Therefore,

25 even though the bulk concentration of the analyte
in the fluid being measured may be low, the an-
alyte concentration at the electrode surface will be
relatively high, thereby permitting detection by vol-
tametry. The irreversibility of antibody-antigen

30 binding reduces the minimum detectable analyte
concentration significantly below that required in
other voltametric measurement applications, such
as measuring trace organic in a plating bath analy-
sis system. Another factor contributing to the high

35 sensitivity is the second harmonic filtering effect
described above.

The sensitivity of the double layer capacitance
to changes in analyte concentration can be in*
creased by adding another species to the fluid

40 containing the analyte. The added species is pref-
erably one which adsorbs at the electrode surface
when the complexed analyte desorbs, and vice
versa. In this way, the double layer capacitance is
affected not only by desorption of the complex, but

45 also by its replacement with the new species. The
added species may have a charge if the antibody-
analyte complex is essentially neutral If the com-
plex is highly charged, the added species may
either be neutral or have a charge opposite to that

50 of the complex. This wilt tend to Increase the
change in the double layer capacitance during de-
sorption of the complex and adsorption of the ad-
ded species as well as during adsorption of the
complex and desorption of the added species.

55 An exemplary suitable additional species is a
surfactant. A surfactant has the desirable property
of adsorbing at interfaces such as that provided by
an electrode. The adsorption of a surfactant is also

6

11 EP 0 640 832 A2 12

less dependent upon charge than the adsorption of
other substances. Surfactants with no charge or
with either positive or negative charge can easily
adsorb. The advantage of using a surfactant is that
the choice of charge or lack of charge can be
made without regard to whether the species will
adsorb. Surfactants of many different types are
commercially available, and specialized surfactants
can be synthesized. Other species capable of ad-
sorption could also be used, including but not
limited to proteins and salts. The choice of material
for the added species will typically depend upon
the characteristics of the analyte being measured.

In a plot of response current magnitude versus
dc sweep voltage, the response current peak due
to adsorption of antibody-antigen complexes will
appear at a different voltage than the peak due to
the adsorption of uncomplexed antibodies alone.
As discussed above, the magnitude of the current
peak corresponding to the antibody-antigen com-
plex adsorption is indicative of antigen concentra-
tion level. However, as the peak representing the
complex grows in magnitude, the peak represent-
ing the pure antibody will typically get smaller In
magnitude. It therefore may be useful in certain
applications to monitor the reduction in antibody
peak height instead of or in addition to monitoring
the complex peak height. For example, the dc
sweep voltage at which the complex peak is ob-
servable might also be the voltage at which an
undesirable electrochemical reaction occurs at the
electrode. This reaction may destroy the antibody-
antigen complex or cause additives or impurities in
the fluid being measured to interfere with the pro-
cesses occurring at the electrode surface. One
such additive may be the surfactant added to the
fluid to enhance the magnitude of the changes in
double layer capacitance as discussed al)ove. With
certain types of additives it may thus become
necessary to measure reduction in a pure antibody
response current peak instead of or in addition to
the antibody-antigen complex response current
peak. The use of both peaks also provides a re-
liability check. For example, a reduction in the pure
antibody peak wrthout a corresponding appearance
of or increase in a complex peak may indicate an
interference with the normal electrode process.

Although the above description is directed to
detection of changes in the double layer capaci-
tance following adsorption and desorption of com-
plexes, other types of reactions may also produce
detectable changes in the double layer capaci-
tance. For example, certain analytes may be sub-
ject to oxidation and reduction during ac voltame-
try. Once such an analyte Is complexed. the oxida-
tion and reduction process can also affect the
double layer capacitance. It is therefore possible to
use double layer capacitance measurements to de-

tect other reactions such as oxidation and reduction
in order to measure concentration of certain an-
alytes.

In order to optimize the accuracy of the re-

5 sponse current spectra produced in accordance
with the preferred ac voltametric technique de-
scribed above, a number of independent physical
test parameters may be varied. These parameters
include, but are not limited to. ac signal amplitude

10 and frequency, dc sweep signal voltage range and
sweep range rate, and ac response signal harmonic
measured. These parameters should be varied in-
dependently to determine the optima! parameter for
use in a particular application. In general, certain

75 settings of the above physical test parameters are
particularly well-suited for monitoring analyte con-
centration in accordance with the preferred em-
bodiment of FIG. 1 and the exemplary method
described above. The preferred ac waveform is a

20 sinusoid with an appropriate amplitude and a fre-
quency of about 10 to 1000 Hz. The amplitude
chosen will depend upon the chemical makeup of
the fluid being tested. An appropriate ac signal
amplitude should provide measurable changes in

25 the response current while avoiding undesirable
reactions between the electrode and the fluid con-
stituents. The prefen'ed frequency range is de-
signed to maximize the quadrature component of
the response current for typical applications. Maxi-

30 mizing the quadrature current will maximize the
sensitivity to changes in the electrode double layer
capacitance. At very high ac signal frequencies, the
electrode double layer capacitance will effectively
become a short circuit and thereby cause the

35 quadrature component of the current to go to zero.
At very low ac signal frequencies, the capacitance
behaves like an open circuit, also reducing the
quadrature component of the response current to
zero. It should be understood that in certain ap-

40 plications ac signal frequencies outside the above
preferred range may be desirable. The dc sweep
signal is preferably swept over an amplitude range
which encompasses adsorption and desorption vol-
tages for the antibodies and antibody-antigen com-

45 plexes. A preferred sweep rate will typically be an
order of magnitude below the ac signal frequency.
As previously described, optimal spectral peak res-
olution is usually obtained using the quadrature
component of the ac response current second har-

50 monic.

An altemative method for use with the elec-
trochemical immunosensor system of the present
invention detects a binding event by measuring its
effect on the impedance at the sensing electrode

55 10. This method is based upon the formation of an
ion gate at the electrode surface. An ion gate is a
channel for migration and diffusion of ions from the
electrode surface to the fluid under test. The chan-

7

13

EP 0 640 832 A2

14

nel can be closed when an antibody on the elec-
trode surface binds an antigen in the fluid. Since
the ion gate controls the flow of a relatively large
quantity of Ions between the fluid and the elec-
trode, a single binding event can produce a mea-
surable change in the response current. The Ion
gate can therefore be considered the chemical
analogue of a transistor. An ion gate is formed by
attaching antibodies to gold particles, or other bind-
ing agents wtiich are themselves attached to the
electrode surface.

In accordance with this alternative method an
indicator ion is added to the fluid sample being
measured so that an ionic response cun'ent will
flow through the sensor. The indicator ion need not
be capable of reacting electrochemically with the
electrode. Any type of ion which does not other-
wise unfavorably interact with the system may be
used. Preferred Indicator ions include chloride salts
such as sodium chloride and potassium chloride or
other types of salts. Salts are preferred because
they tend to maintain a neutral pH and are there-
fore less likely to affect the pH sensitive proteins
used as binding agents or analytes. An ionic re-
sponse current is then developed when an ac vol-
tametric signal of sufficient frequency is applied ot
the potentiostat. An ac signal frequency of about
1,000 to 10,000 Hz is suitable for maintaining an
ionic response current despite the lack of an elec-
trochemical reaction between the indicator tons and
the electrode. The ionic current flow can be main-
tained via charging and discharging of the elec-
trochemical double layer capacitance at the elec-
trode. The formation of an antibody-antigen com-
plex closes an ion gate and results in a measurable
decrease in the ionic response current flow.

Electroactive indicator ions capable of elec-
trodeposition onto the electrode surface may also
be used. One exemplary electroactive indicator ion
is nickel. In the case of electroactive indicator ions,
an ionic response current is developed as a result
of the electroactive indicator ions being plated to
and stripped from the electrode surface in re-
sponse to the applied voltametric signal. A binding
event can then be readily detected as a decrease
in the ionic response current normally associated
with plating and stripping of the electroactive in-
dicator ions.

A preferred technique for forming the ion gate
is to adsorb antibodies onto colloidal gold particles
which include previously adsorbed protein A or G,
and then apply the resulting colloidal gold-antibody
complex to the surface of the gold electrode using
electrophoresis or any other suitable process. A
gold electrode Is preferred for the reasons dis-
cussed above in the general description of the
electrochemical immunosensor system. Colloidal
gold with adsorbed protein A or G is a commer-

cially available product. The proteins A or G are
preferred because they bind to the Fc region or
inactive end of an antibody. This allows the active
end of the antibody to extend into solution. The
5 colloidal gold-antibody complex will adsorb to the
gold electrode, with the antibody acting as a glue
between the colloidal gold particle and the gold
electrode.

The colloidal gold particles are generally
10 spherical in shape, and as they accumulate on the
electrode surface, they will close pack together,
leaving small pores between them. The antibodies
extend from the surface of the particles into the
pore, such that they are exposed to antigens within
75 the fluid. When an antibody binds an antigen, the
antigen plugs the pore and thereby closes the ion
gate, preventing further passage of ions through
the gate. Colloidal gold is available in particle sizes
as small as 5 nanometers, and the pores between
20 colloidal gold-antibody complexes will therefore be
readily plugged by most antibody-antigen com-
plexes.

The operation of the altemative method is as
follows. An electrical signal from waveform gener-

25 ator 5 is applied to potentiostat 8 to which a
sensing electrode 10 has been connected as de-
scribed earlier and as shown in FIG. 1. In this
method the ac electrical signal preferably consists
of an ac voltametric signal without any dc sweep

30 component. A constant dc potential offset may be
added to the ac signal in order to avoid undesirable
parasitic reactions. Although it is preferred to keep
the dc offset around zero to avoid most parasitic
reactions, it may be beneficial to add a positive or

35 negative offset in certain applications. The sensing
electrode 10 has a quantity of ion gates formed on
it surface by closely packed colloidal gold-antibody
complexes adsorbed or otherwise affixed to its
surface. A quantity of indicator ions have been

40 added to the fluid containing the analyte to be
measured as described above. The ionic response
current developed through sensor 9 is then mon-
itored using the lock-in amplifier 6 and strip chart
recorder 7 in the manner described previously. The

45 first harmonic of the in-phase component of the ac
response current is preferably monitored. In this
method, the peak seen on the strip chart recorder
7 will not occur at a particular voltage determined
by a sweep signal potential as in the first method

50 described, but instead will occur as soon as the
electrode is exposed to the analyte and complex-
ation begins. The magnitude of the ionic response
current provides a sensitive indication of analyte
concentration.

55 The accuracy of the response current as an
indicator of particular analyte concentrations can be
optimized by independently varying system param-
eters including ac signal amplitude and frequency,

8

15 EP 0 640 832 A2 16

constant dc offset and ac harmonic measured. The
system parameters preferred for many applications
include an ac signal with a frequency of about
1,000 to 10,000 Hz. a dc offset of zero volts, and
measurement of the first harmonic of the response
current. The ac signal amplitude will again depend
upon the binding agent, analyte and other constitu-
ents within the fluid under test. The amplitude is
chosen such that response current resolution is
maximised without causing undesired electro-
chemical reactions at the electrode. Use of the first
harmonic will filter out the constant background
current flowing through those ion gates which have
not been closed by complexation. The first har-
monic can provide this filtering function since the
applied voltametric signal no longer includes a dc
voltage sweep as in the first method.

FIG. 2 shows an alternative set of measure-
ment equipment for use in conjunction with the
above-described ion gate method. The waveform
generator 5, lock-in amplifier 6. strip-chart recorder
7 and potentiostat 8 of FIG. 1 can be replaced by
the ac impedance bridge 35 of FIG. 2. The ac
impedance bridge 35 measures the sensor imped-
ance by applying an ac signal and monitoring the
current through the sensor. The ac impedance
bridge 35 can thus serve as both a signal generat-
ing means and a signal monitoring means.

The impedance bridge 35 generates an ac
signal of appropriate amplitude and frequency and
applies it via leads 36 and 37 to sensor 39. The
sensor 39 includes a sensing electrode 40 and
counter electrode 42. When an ac impedance
bridge is used as a signal generating and monitor-
ing means a reference electrode need not be in-
cluded in sensor 39. Sensor 39 contains a quantity
of indicator ions and sensing electrode 40 has a
quantity of antibodies present on its surface in
accordance with the above described ion gate
method. The applied ac signal from impedance
bridge 35 causes an ionic response current to flow
between electrodes 40 and 42. Antibodies com-
plexing with analytes within sensor 39 close ion
gates on the sensing electrode surface and thereby
increase the impedance and reduce the ionic re-
sponse current flowing through sensor 39. The
increased impedance Is monitored by impedance
bridge 35 and is proportional to the number of
complexes formed at the sensing electrode sur-
face, which is a function of analyte concentration.

Commercially available impedance bridges
typically provide an ac signal at about 1.000 Hz.
This is within the frequency range discussed at>ove
which permits an Ionic current resulting from charg-
ing and discharging of the electrode double layer
capacitance. The indicator ion within sensor 39
therefore need not be electroactive. Although under
certain circumstances the 1,000 Hz typically pro-

vided by commercially available impedance brid-
ges may be too low a frequency for optimal opera-
tion, custom designed impedance bridges could be
constructed to provide a desired frequency of op-

5 oration.

Although the above description has been di-
rected to the use of exemplary voltametric tech-
niques to detect antibody-antigen binding events
and thereby antigen concentration levels, this is by

TO way of illustration and not limitation. The elec-
trochemical immunosensor system and methods of
the present invention can be applied to detect any
analyte which binds with any binding agent to form
a complex. Other voltametric techniques and

75 equipment could also be applied to analyte detec-
tion using the methods of the present invention. It
will be understood by those skilled in the art that
many alternate embodiments of this invention are
possible without deviating from the scope of the

20 invention, which is limited only by the appended
claims.

Claims

25 1. An electrochemical immunosensor system
adapted for use in measuring a concentration
of analytes in a fluid, said system comprising:

an immunosensor having at least one
sensing electrode for contact with said fluid,

30 said sensing electrode having a surface on

which a binding agent is present, such that
said binding agent binds a portion of said
analyte in said fluid to form a number of com-
plexes;

35 signal generating means for generating an

electrical signal for application to said sensing
electrode, said electrical signal producing a
response current through said sensing elec-
trode, said response cunrent having at least

40 one signal characteristic which is dependent
upon said number of complexes; and

signal monitoring means for monitoring
said signal characteristic of said response cur-
rent.

45

2. The immunosensor system of claim 1 wherein
said signal generating means includes:

a waveform generator which provides an
ac signal; and

50 a potentiostat having an input to which

said ac signal is applied and an output elec-
trically connected to said sensing electrode.

3. The immunosensor system of claim 2 wherein
55 said signal generating means further includes a

dc signal generator which provides a dc sweep
signal, and further wherein said electrical sig-
nal applied to said sensing electrode through

9

17

EP 0 640 832 A2

18

said potentiostat includes said ac signal super-
imposed on said dc signal.

4. The immunosensor system of claim 1 wherein
said signal monitoring means includes phase-
locking means for phase-locking said response
current to said etectricai signal.

5. A method of measuring an analyte concentra-
tion using an immunosensor, said method
comprising the steps of:

providing a fluid containing an analyte con-
centration to be measured;

providing an immunosensor having at least
one sensing electrode, said sensing electrode
having a surface on which a binding agent is
present;

placing said sensing electrode of said im-
munosensor in contact with said fluid such that
a portion of said binding agent on said sensing
electrode surface binds a portion of said an-
alyte in said fluid to form a number of com-
plexes;

applying an electrical signal to said sens-
ing electrode on which said complexes are
formed, said electrical signal producing a re-
sponse current flowing through said sensing
electrode having at least one signal character-
istic which is dependent upon said number of
complexes; and

monitoring said signal characteristic of
said response current.

6. The method of claim 5 further including the
step of adding an additional species to said
fluid containing said analyte, said additional
species capable of adsorbing to said electrode
at a voltage at which said complex desorbs
from said electrode and desorbing from said
electrode at a voltage at which said complex
adsorbs to said electrode.

7. The method of claim 5 wherein said electrical
signal is a voltammetric signal including an ac
signal superimposed on a dc sweep signal,
said ac signal having an amplitude and a fre-
quency.

8. The method of claim 7 wherein said step of
monitoring said signal characteristic of said
response current includes measuring an ac
component of said response current as said dc
signal is swept over a potential range, said
measurement of said ac component of said
response current in relation to said dc potential,
range being expressed as ac current spectra.

9. The method of claim 8 wherein measurement
of said ac current is made at the second
harmonic frequency relative to the frequency
of said ac signal.

5

10. The method of claim 5 wherein said fluid fur-
ther contains a quantity of indicator ions and
wherein said binding agent on said surface of
said electrode forms a quantity of ion gates on

10 said surtace, such that when said binding
agent binds said analyte in said fluid to pro-
duce a complex, said complex closes one of
said ion gates and thereby reduces a mag-
nitude of said response current.

75

10

EP 0 640 832 A2

FIG. 1.

11

EP 0 640 832 A2

35

39

FIG. 2.

12